Magnetic resonance imaging involves magnetically nutating nuclear spins of an object placed in a static magnetic field with an RF signal at a Larmor frequency. MR signals emanate therefrom as the nutated spins relax and an image is generated based on the received MR signals. Since large pulses of electrical current are supplied in an MRI pulse sequence to a gradient magnetic field coil in the static magnetic field, mechanical distortion of the gradient magnetic field coil occurs. Such mechanical distortion can generate loud noises during imaging and this can be painful to a patient. However, in recent years, silencing technology has been developed where the gradient magnetic field coil is surrounded by a vacuum vessel and the noise problem is reduced.
In order to perform spatial encoding of the MR signals, an MRI pulse sequence of currents whose rise and fall time is typically several 100 micro seconds and whose magnitude can be about 100 A–200 A, is supplied to the gradient magnetic field coil. High rise/fall pulse speeds and large amplitude current pulses also generate high heat in the coil structures. In order to dissipate (i.e., remove) the accumulated heat, air cooling and water cooling methods can be used together. However, water cooling alone may have to be used for cooling a gradient magnetic field coil that includes a noise-reducing vacuum vessel, because it is difficult to use air cooling in the vacuum vessel.
Recently, high speed imaging methods, such as echo planar imaging (EPI), etc., have been put into practical use by improving the gradient magnetic field power supply. In such high speed imaging methods, due to high power gradient magnetic fields and high speed switching, fast rise/fall large magnitude pulse currents are required. Leakage magnetic fields also increase according to the pulse current magnitude. Generally, in order to shield spurious or leakage magnetic field flux from the imaged volume, active or passive magnetic shielding is used. With active shielding, a reverse current is supplied to an actively shielded gradient coil (ASGC) disposed near the gradient magnetic field coil. With passive shielding, an iron base shield material is typically used.
However, driving currents generated in these gradient magnetic field coils and shielding coils also generate heat. The shielding material itself makes the empty space narrower, which makes it more difficult to dissipate (i.e., remove) the heat. In order to decrease heat generation, it is generally effective to use larger size coil wire to reduce its resistance, or to enlarge a cooling water pipe cross section to increase the speed of flowing coolant. However, since there is only a relatively narrow empty space near the gradient magnetic field coil, it is also difficult to enlarge the coil wire size and/or the cold water pipe size. Thus, cooling by such hardware methods is limited.
If the available hardware cooling is not sufficient for all conditions, it is desirable to observe the gradient magnetic field coil temperature, and to stop imaging if the coil temperature rises beyond an allowable maximum value. Conventionally, a coil temperature sensor is attached on the surface of resin surrounding the gradient magnetic field coil. However, since the heat conductivity of the resin is relatively low, it takes some dozens of seconds to measure actual coil temperature changes by such a temperature sensor. That is, the MRI apparatus may be damaged because of the time-delayed response of the coil temperature sensor.
There are also prior art MRI systems which calculate coil temperature based on the current supplied to the gradient magnetic field coil. For example, see Japanese patent documents 10-71131; 10-43156; 10-43157; 10-71132 and/or 8-592.